1. Field of the Invention
The present invention relates generally to the fields of chemistry, biology, biochemistry, medical devices and chemical detection devices. More particularly, it concerns polymeric compositions, methods and devices for the detection of one or more selected analyte(s), preferably detection of one or more analyte(s) in vivo. The present invention also concerns polymeric constructs for analyte detection and concentration measurement using visualization or production of electrochemical signals from a construct after implantation in vivo. In particular, the invention concerns detection or measurement of blood glucose levels in vivo for the management of diabetes. Additionally, devices are provided to detect and measure the optical or electrochemical changes in polymeric constructs.
2. Description of Related Art
Diagnosis of various disease and injury conditions are often made on the basis of the detection and measurement of the concentration of one or more body chemicals or analytes. Currently, the levels of numerous selected body chemicals or analytes are measured manually and invasively by withdrawing a sample of blood. For most analytes the sample is usually sent to a centrally located lab where it is typically analyzed with large and expensive machines using wet chemistry, immunoassays, and/or enzyme electrode based biosensing. This mode of operation is expensive and time consuming, and therefore represents a significant hazard to critically ill patients in operating rooms, intensive care units and trauma/critical care units. In particular, for trauma/critical care cases it is known that survival rates decrease dramatically if treatment is delayed for more than one hour. For instance, the measurement of arterial blood gases is a primary indicator of respiratory function, and lactate values are used as an indicator of shock. Therefore, the frequent assessment of these and other analytes is essential to clinical diagnosis and management.
Exemplary of diseases that require frequent monitoring of analytes is diabetes mellitus. Diabetes mellitus is a chronic disease that, if unregulated, can give rise to large fluctuations in blood glucose levels. This disease currently afflicts over 100 million people worldwide and nearly 14 million in the United States (National Institute of Diabetes and Kidney Diseases, 1994). In the U.S. this disorder, along with its associated complications, is ranked as the seventh leading cause of death (Cotran et al., 1989). In order to maintain normal glucose levels, blood glucose must be monitored frequently throughout the day. Self-monitoring of blood glucose is recommended for diabetic patients as the current standard of care and, since the announcement of the Diabetes Control and Complications Trial results (National Institute of Diabetes and Digestive and Kidney Diseases, 1993), there is now no question that intensive management of blood sugars is an effective means to prevent or at least slow the progression of diabetic complications such as kidney failure, heart disease, gangrene, and blindness (National Institute of Diabetes and Digestive and Kidney Diseases, 1993; Wysocki, 1989; Speicher, 1991).
The goal of diabetes therapy is to approximate the 24-hour blood glucose profile of a normal individual. Without regulation, hypoglycemia, a condition in which the blood glucose level falls well below normal, can result, which can cause the patient to slip into a coma and eventual death. Alternatively, a condition known as hyperglycemia can develop, in which blood glucose levels can rise considerably above normal levels. If left untreated, these abnormally high blood glucose levels may result in long-term complications such as an increased risk of coronary artery disease, hypertension, retinopathy, neuropathy, and nephropathy (National Institute of Diabetes and Kidney Diseases, 1994; Cotran et al., 1989; National Institute of Diabetes and Digestive and Kidney Diseases, 1993; Hanssen, 1986).
Proper treatment includes maintaining blood glucose levels near normal levels. This can only be achieved with frequent blood glucose monitoring so that appropriate actions can be taken, such as insulin injections, proper diet, or exercise. Unfortunately, the currently preferred method of sensing is an invasive technique, requiring a finger stick to draw blood each time a reading is needed. This approach is both time-consuming and painful. Therefore, there is a lack of compliance among the diabetic population for even monitoring their levels once per day, not to mention the recommended five or more times daily (National Institute of Diabetes and Digestive and Kidney Diseases, 1993).
One potential method of achieving tighter metabolic control in diabetic patients is a closed-loop insulin delivery system, incorporating a microprocessor-controlled insulin pump and a glucose sensor. Various amperometric devices have been fabricated based upon the electrochemical oxidation of H2O2 generated during a reaction between glucose and oxygen catalyzed by glucose oxidase (Tatsuma et al., 1994). The focus of amperometric biosensors appears to have shifted towards the incorporation of charge mediators as electron xe2x80x9cshuttlesxe2x80x9d between the redox center of the enzyme and the electrode surface (Hale et al., 1991; Pishko et al., 1991). These H2O2 and mediator based biosensors have taken on many forms, including covalent immobilization directly to the electrode surface, retention by a membrane, or entrapment in a polymer hydrogel (Henning, T. and Cunningham, D., 1998).
Glucose sensors, which use an enzyme (glucose oxidase) to achieve specificity, are currently not stable or sensitive enough to meet the demands of a closed-loop delivery system. As a result, the application of glucose biosensors has been primarily limited to home glucose test meter and blood-gas instruments containing sensors for glucose (Rouhi, 1997). There are a number of reasons for this lack of commercial progress, both technical and economic. Technically, many proposed biosensors for glucose simply do not have the accuracy and stability (operational or storage) to meet the desired need. Inaccuracy and imprecision in sensor performance are frequently due to imprecision in sensor manufacturing, e.g. immobilized biomolecules cannot be deposited on transducer surfaces at the same density and with the same mass transfer limitations. Instability is often a problem inherent in the biomolecule, the result of poor immobilization methods resulting in leaching, or inactivation of the biomolecule by species present in the sensing environment (Pishko, M. V., 1995).
Glucose detection devices have been reported that quantify glucose concentration in blood and body fluids. One such device uses fluorescence resonance energy transfer (FRET) between a labelled ligand and a labelled carbohydrate-containing receptor (U.S. Pat. No. 5,342,789). The binding of glucose to the receptor prevents energy transfer from the labelled receptor to the labelled ligand, and thus prevents the quenching of the flourescence of the labelled receptor.
Noninvasive methods to quantify blood chemicals, particularly glucose, have been attempted using various optical approaches. Four primary approaches being investigated, including near-infrared (NIR) absorption spectroscopy (Small et al., 1993; Marbach et al., 1993; Robinson et al., 1992; McShane et al., 1997), NIR scattering (Kohl and Cope, 1994; Maier et al., 1994), polarimetry (March et al., 1982; Gough, 1982; Cameron and Cotxc3xa9, 1997; King et al., 1994; Cotxc3xa9 et al., 1992), and Raman spectroscopy (Goetz, Jr. et al., 1995; Berger et al., 1997).
Each of these approaches suffer primarily from a lack of specificity. The NIR scatter approach is confounded by changes in indices of refraction, since tissue scattering is also caused by a variety of substances and organelles which all have different refractive indices. The Raman approach is non-specific, lacks good sensitivity, requires high powers, and suffers from large background autofluorescence of the tissue in vivo (Cotxc3xa9, G. L., 1997).
Although reports have been given for in vivo near-infrared (NIR) absorption spectroscopy data, these results have been primarily based on bolus injections of intravenous glucose (Marbach, 1993; Robinson 1992). Since the multivariate statistics can produce strong temporal correlation due to independent factors these results are suspect and could likely be due to temporal variations other than glucose. Another drawback to the NIR approach is a lack of repeatability of the NIR signal in vivo both within and especially between patients (Day, 1996; Sabatini, 1996). Thus, the main drawback for the NIR technique is again the lack of specificity as well as of lack of sensitivity in the presence of these confounders.
For polarimetry to be used as a noninvasive technique for blood glucose monitoring, the signal must be able to pass from the source, through the body, and to a detector without total depolarization of the beam. Both the skin and eye have been used as detection sites for this technique. Overall, tissue birefringence and motion artifact are sources of error for this approach regardless of the sensing site. The change in rotation due to other chiral molecules such as proteins creates problems of specificity. Thus, as with all the previously described potentially noninvasive optically based approaches the primary drawback is not only the lack of sensitivity but the lack of specificity of these approaches.
The use of fluorescence was investigated for glucose monitoring in vivo using an indwelling fiber optic approach and a membrane (Mansouri and Schultz, 1984; Schultz et al., 1982), however, to date it has not been developed as a noninvasive technique. In this approach the fluorophore was bound on the inner surface of the membrane at the tip of the fiber. Glucose had a higher affinity for the membrane base molecule and displaced the fluorophore causing the fluorescent light to be returned through the fiber and thus had good specificity. However, the indwelling optical fiber-based approach has many of the same problems associated with previous electrochemical approaches, including membrane fouling, encapsulation, and decrease in response time, as well as opening the body to potential infection. In addition, the fiber was intravenously implanted and thus was small, yielding very low signal-to-noise-ratios (Mansouri and Schultz, 1984; Schultz et al., 1982).
Biosensor applications for the detection and identification of pathogens by DNA or RNA hybridization, or rapid DNA sequencing require a high-density pattern of individual sensing elements (Chee et al., 1996; Yershov et al., 1996). Thick film technology have been used for a number of years to fabricate single biosensors for the home glucose test market, but this technology is not amenable to the fabrication of micrometer scale arrays of sensors. Biotin/Avidin systems have been utilized to immobilize enzymes in an ordered fashion, albeit without charge mediators (Dontha et al., 1997). If biosensors are to see more wide spread application, sensor fabrication technologies must be developed that allow the development of stable, easily manufactured multisensor arrays (Madou and Tierney, 1993).
Recent research on patterning biomolecules on surfaces has focused primarily on self-assembled monolayers (SAMs; Mooney et al., 1996; Whitesides et al., 1991), and tethered biomolecules on surfaces that may potentially form addressable patterned arrays (Mooney et al., 1996; Britland et al., 1992; McLean et al., 1993). These patterned surfaces are formed, particularly using alkane thiols and their derivatives on gold-coated surfaces. SAMs also permit the site specific immobilization and orientation of biomolecules on a surface. However, two-dimensional approaches such as SAMs may limit the number of biomolecule recognition sites on the sensor surface, and thus may have low signal levels and require shielding or other measures to reduce noise. The structure of self-assembled molecules on a surface can also result in defects or xe2x80x9cpinholesxe2x80x9d in the monolayer and contribute to instability, particularly at applied potentials. The current adhesion chemistry used in the fabrication of SAMs also permits monolayer formation only on a limited number of surfaces, most commonly gold. In addition to monolayers, photolithography and other photoinduced patterning chemistries were highlighted in a few studies, demonstrating the formation of patterned biomolecule surfaces (Sundberg et al., 1995; Dontha et al., 1997) and micropatterned polymers for optical chemical sensing (Healey et al., 1995).
In prior research of biosensors based upon redox polymers coupled to biorecognition molecules such as oxidoreductases, the polymer served to immobilize the enzyme via formation of an insoluble protein/polymer complex (Pishko et al., 1990a; Pishko et al., 1990b; Tatsuma et al., 1994), through the physical entrapment of the enzyme in a polymer film (Hale et al., 1990; 1991), and/or through the covalent cross-linking of the enzyme and polymer (Gregg and Heller, 1991b; Ohara et al., 1993). Amperometric biosensors based on redox polymer/enzyme complexes were shown to be miniaturizable (Pishko et al., 1991) and could measure analytes either intravenously or subcutaneously when implanted in rats (Csoregi et al., 1994; Kerner et al., 1993; Linke et al., 1994; Quinn et al., 1995b; Schmidtke et al., 1996). In all of these studies, the redox polymers were synthesized by heat-induced free radical polymerization. Electropolymerization of vinyl-containing redox monomers was used to polymerize and deposit redox polymers on electrodes but was not demonstrated as an effective method of immobilizing enzymes or forming patterned films (Denisevich et al., 1982; Abruna et al., 1981). The photopolymerization of redox polymers was previously reported for vinylferrocene/acrylamide copolymers (Nakayama et al., 1995).
One technique in biosensor construction involves the building of individual monolayers on surfaces based upon the attraction between oppositely charged species. Issues involved in the development of patterned polyion multilayers that have been examined include solution ionic strength and number of multilayers (Hammond and Whitesides, 1995; Clark and Hammond, 1998; Gregoriou et al., 1997). Typically, these systems utilized Sulfonated Polystyrene as the polyanionic component and compounds such as Poly(allylamine hydrochloride) as the polycationic component. These polyion multilayers were grown in distinct patterns through the use of micro-contact printing and blocking agents anchored to a gold substrate. The effect of solution ionic strength on these multilayers has been examined (Clark et al., 1997; Sukhorukov et al., 1996) as well as the attachment of redox-active osmium complexes via similar techniques (Bretz and Abruna, 1996). Alternating glucose oxidase (GOX) and charge mediator layers for the fabrication of a glucose enzyme electrode have been utilized (Chen et al., 1969; Hou et al., 1997; Hou et al., 1998). Other groups have also performed similar work utilizing a ferrocene derivative as their poly-cationic charge mediator (Hodak et al., 1997).
PEG-based coatings have been used to improve the biocompatibility of implanted glucose sensors and demonstrated that these hydrogels were not glucose mass transfer limiting (Quinn et al., 1995). PEG-based polymers have previously been evaluated for in vivo use as protein drug delivery devices, for postoperative adhesion prevention, and for biocompatible membranes over electrochemical sensors (West and Hubbell, 1995; Pathak et al., 1992; Sawhney et al., 1994). The stability and solubility of numerous proteins, including bovine serum albumin, catalase, and interleukin-2, is reportedly increased upon conjugation to PEG (Delgado et al., 1992). Monomethoxy poly(ethylene glycol)-5000 has been conjugated to Con A while retaining Con A""s sugar binding abilities (Mattiasson and Ling, 1980). Lakowicz and co-workers (Lakowicz and Maliwal, 1993) have also developed fluorescent assays for glucose, based on phase-modulation fluorimetry and Con A/dextran moieties. These studies were conducted in an aqueous solution, and it was indicated that a polymeric acceptor may be used to shield the glucose sensor behind a glucose-permeable barrier (Lakowicz and Maliwal, 1993).
Work invoking PEG and glucose oxidase focused on pH-sensitive hydrogels which swelled and shrank as a result of glucose concentration. The large physical changes needed to measure glucose concentrations were slow to develop, and limited due to the swelling resulting in an influx of buffered solution, which reduced or eliminated the small change in pH (Hassen et al., 1997).
Despite these attempts at alternative methods of analyte detection, there is still a need to design and develop technology that would provide an either minimally invasive or noninvasive method to measure biological analytes or external chemicals. There is a need for devices and methods to easily allow an increase in the accuracy and frequency of measurement, identify potentially hazardous compounds, provide for tighter control of patient compliance, while fostering fewer secondary complications than current methods. Such devices and methods would thus represent a significant advance in the art.
The present invention overcomes these and other deficiencies present in the art by providing a variety of chemically sensitive, stable (insoluble over a specified period of time), nontoxic, and non-antigenic hydrogel structures, wherein the structure may be, for example, in the form of a particle or in the form of a hydrogel adherent to a substrate, such as an electrode substrate. In preferred embodiments, two or more electodes are used. However, electrodes that are not in contact with the hydrogel may also be used in certain embodiments. The hydrogel structure may undergo a measurable change in at least one electrochemical or optical property as a function of interaction with one or more substance(s) to be detected. Also provided are methods of using these hydrogel particles to detect one or more selected analytes, either in in vivo or external to a living organism. The present invention provides an implantable non-invasive monitoring approach, which may provide better compliance and reduce the risk of infection. Additionally provided are devices used to detect and measure the optical or electrochemical changes.
The present invention first provides a composition comprising an analyte sensitive compound comprised within a hydrogel, wherein the hydrogel comprises a polymerized material including but not limited to poly(ethylene glycol), poly(ethylene glycol)-co-anhydride, poly(ethylene glycol)-co-lactide, poly(ethylene glycol)-co-glycolide, poly(ethylene glycol)-co-orthoester, poly isopropylacrylamide, polyHEMA, polyacrylamide, sodium alginate or a combination thereof. In certain embodiments, the hydrogel is a macromer, or construct made of polymerized material or polymers that are themselves polymerized by covalent attachment at the ends of the individual polymers that comprise the macromer. Polymerized material or polymers that may be polymerized in the macromer, include but are not limited to, poly(ethylene glycol), poly(ethylene glycol)-co-anhydride, poly(ethylene glycol)-co-lactide, poly(ethylene glycol)-co-glycolide, poly(ethylene glycol)-co-orthoester, poly isopropylacrylamide, polyHEMA, polyacrylamide, sodium alginate or a combination thereof. As defined herein, a xe2x80x9chydrogelxe2x80x9d refers to a polymeric material that allows a fluid or aqueous medium to diffuse throughout the material. This property of rapid diffusion allows rapid contact of the hydrogel and its components with substances dissolved or dispersed within the fluid or aqueous medium. The hydrogel may be polymerized by any technique known to those of ordinary skill in the art, such as for example, chemical induced polymerization or photopolymerization. The hydrogel may act as a support or framework for additional components, such as analyte sensitive compounds, monomers, comonomers, and other materials. The hydrogel support may covalently bind or non-convalently entrap the additional materials.
In various embodiments of the present invention, the composition may further comprise additional monomers to alter the properties of the hydrogel in advantageous ways. A monomer that is contemplated as being useful in the present invention include, but is not limited to, a vinyl-containing monomer, an acrylate-containing monomer, a methacrylate-containing monomer or a combination thereof. The hydrogel may be copolymerized with a monomer, including but not limited to, a vinyl-containing monomer, acrylate-containing monomer, ethacrylate-containing monomer or combination thereof. A monomer may function to cross-link the hydrogel, or foster contact and/or binding of multiple hydrogels. Such contact may be fostered by cross-linking the plurality of hydrogels with the monomers and/or enhancing the affinity of an at least first body of hydrogel material for an at least second body of hydrogel material by conferring opposite charges to each successive body or layer of hydrogel material. Additionally, a monomer may have specific affinity for, or be capable of chemically linking to, an analyte sensitive compound or additional material.
In certain aspects, a vinyl-containing monomer is acrylic acid, allyl amine, styrene, allyl alcohol, acrylamide, acrylate-PEG-hydroxysuccinimde ester, Os(vinyl pyridine)(bis-bipyridine)2Cl, vinyl imidazole, vinyl bipyridine, vinyl ferrocene, styrene, pentadiene, methyl pentadiene or polyacrylated monomer. In other aspects, one or more monomers comprise the composition. As used herein certain embodiments, a comonomer is two or more monomers used to comprise the composition. In certain embodiments, the hydrogel comprises a copolymer of two or more polymers. The copolymer may further comprise a monomer or comonomers. The monomer or comonomer may include, but is not limited to, a hydrophobic, a cationic, an anionic, a neutral but hydrophilic, a biomolecule reactive, a redox, or a multifunctional crosslinking monomer or comonomer, or a combination thereof. As used herein certain embodiments, a biomolecule reactive monomer or comonomer contains a reactive group such a vinyl, methylacrylate or acrylate moiety to covalently bond to the hydrogel composition, as well as being able to bind to a biomolecule, including but not limited to, proteins, nucleic acids or lipids. As used herein certain embodiments, a redox reactive monomer or comonomer contains a chemical group such a vinyl, methylacrylate or acrylate moiety to covalently bond to the hydrogel composition, as well as being able to bind to a redox reactive moiety, i.e. a compound selected for its ability to undergo oxidation or reduction, or oxidize or reduce other compounds. As used herein certain embodiments, a multifunctional monomer or comonomer contains more than one reactive groups such a vinyl, methylacrylate or acrylate moiety to covalently bond to one or more the hydrogel composition(s), as well as being able to bind to other agents, such as, for example, a analyte detection agent. In some embodiments, the hydrogel is copolymerized with one or more comonomers or monomers. A preferred hydrophobic comonomer includes styrene, acrylic acid, methacrylic acid, an alkene, or a combination thereof. A preferred alkene is pentene. A preferred cationic comonomer includes allyl amine or acrylamide. A preferred anionic comonomer is styrene sulphonate. A preferred neutral but hydrophilic comonomer is allyl alcohol. A preferred biomolecule reactive comonomer is acrylate-PEG-hydroxysuccinimide ester, wherein the acrylate moeity is preferred to bind to the hydrogel, and the hydroxysuccinimide is preferred as a biomolecule reactive or binding moiety. A preferred redox comonomer includes vinyl ferrocene, an Os derivative of vinyl pyridine, an Os derivative of vinylimidazole, a Ru derivative of vinyl pyridine or a Ru derivative of vinylimidazole. A preferred multifunctional crosslinking comonomer includes trimethylol propane triacrylate or pentaerythritol triacrylate.
In certain aspects, the hydrogel may further comprise a positively charged polyelectrolyte, a negatively charged polyelectrolyte or a combination thereof.
The analyte sensitive material may be attached to hydrogel or comprised within the hydrogel. In certain preferred aspects, the analyte sensitive compound is non-covalently entrapped in the hydrogel and/or covalently attached to the hydrogel. In a preferred aspect, analyte sensitive compound binds at least one selected analyte. An analyte that may be detected include, but is not limited to, a carbohydrate, a protein, a nucleic acid, a lipid, a chemical (in solid, or preferably, liquid or gas form), or a combination thereof.
The analyte sensitive material may be any compound or grouping of compounds that binds to or interacts with a substance to be detected. The analyte sensitive material may also detect the analyte indirectly by detecting a by-product of its presence, including but not limited to, a chemical degradation product of the analyte or a change in the pH of the medium in contact with the composition. As used herein xe2x80x9cdetect an analytexe2x80x9d or xe2x80x9cdetect a substancexe2x80x9d will be understood to encompass direct detection of the analyte itself or indirect detection of the analyte by detecting its by-product(s). Detection of an analyte may be by contact of the analyte sensitive material with the analyte or its by-product. The analyte sensitive material may be a protein that binds an analyte or its by-product, including but not limited to an enzyme, an antibody, or a lectin. In a preferred aspect, the enzyme is a glucose oxidase or an organophosphate hydrolase enzyme.
In certain embodiments of the present invention, the analyte sensitive compound produces an electrochemical change upon contact with a selected analyte. In a preferred aspect, the composition may further comprise one or more electrodes in operational association with the hydrogel. In a particularly preferred aspect, the electrode detects an electochemical change upon contact of the analyte sensitive compound with a selected analyte.
In certain embodiments of the present invention, the analyte sensitive compound produces an optical change upon contact with a selected analyte. In a preferred aspect, the optical change is a fluorescence change upon contact with the selected analyte. In one aspect, the composition may further comprise an optical detection device in operational association with said hydrogel. In a particularly preferred aspect, the optical detection device comprises an one or more electrodes, transistors, diodes, or photoelectric cells that detects or communicates the detection of an optical change upon contact of said analyte sensitive compound with a selected analyte.
Various mechanisms may be used to produce an optical change upon contact of the analyte sensitive material with the analyte. The analyte may, for example, change the pH of the medium in contact with the analyte sensitive material. The analyte sensitive material may, for example, comprise a pH sensitive dye that undergoes a change in color, fluorescence or phosphorescence upon change of the medium""s pH. In a preferred aspect, the analyte produces a change in pH upon contact with the analyte sensitive material or hydrogel.
In certain aspects, the analyte sensitive compound and/or the hydrogel is in operable association with at least a first fluorescent label. In a preferred aspect the analyte sensitive compound binds an analyte, and the binding of the analyte to the analyte sensitive compound alters, increases and/or reduces the fluorescence of the at least a first fluorescent label. In particularly preferred aspects, the first fluorescent label is HPTS or SNAFL-1.
In certain embodiments, the composition comprises a component that binds to the analyte sensitive compound, wherein binding of the analyte to the analyte sensitive compound decreases the binding of the analyte sensitive compound and the component. In one aspect, the analyte competes with the component for binding to the analyte sensitive compound.
In a preferred aspect, the analyte sensitive compound or construct comprises a first conjugate that produces an optical change or electrochemical change. In another aspect, the analyte sensitive compound or construct comprises a second conjugate that produces an optical change or electrochemical change. The first or second conjugate may comprise a component that binds to the analyte sensitive compound or construct. The component may bind the first or second conjugate. In a preferred aspect, the first and/or second conjugate is a fluorophore conjugate. In another preferred aspect, the fluorescence of the first fluorophore conjugate is quenched by the second fluorophore conjugate. In a particularly preferred aspect, the first conjugate comprises dextran and the second conjugate is concanavalin A. Preferred fluorophores include, but are not limited to, FITC or TRITC. In a particularly preferred aspect, the first fluorophore conjugate comprises FITC and the second fluorophore conjugate comprises TRITC.
An analyte sensitive compound or construct of analyte sensitive compounds for use in the present invention include, but are not limited to, a nucleic acid, a protein, Con A, a Os(vinyl bipyridine)(bis pyridine)2 derivative or Os(vinyl bipyridine)(bis-phenathroline)2. The analyte sensitive protein may bind to a particular analyte, including but not limited to, a lectin that binds to glucose, or an enzyme that binds to a specific analyte, including but not limited to, glucose oxidase, galactose oxidase, cholesterol oxidase, cholesterase, lactate oxidase, glucose dehydrogenase, pyruvate oxidase, lactate dehydrogenase or bilirubin oxidase. In certain preferred aspects, the analyte sensitive compound is an oxidoreductase or a glucose oxidase. In additional aspects of the invention, the analyte sensitive protein is an antibody. In certain preferred aspects, the antibody is a monoclonal antibody. Monoclonal antibodies can be tailor made to preferentially or specifically bind a particular epitope or compound, and thus may be used to grant the same specificity in the detection of a particular analyte in the present invention. The amino acid and nucleotide sequences for various proteins, enzymes and antibodies are well known to those of ordinary skill in the art, and may be found at computerized databases known to those of ordinary skill in the art. One such database is the National Center for Biotechnology Information""s Genbank and GenPept databases (http://www.ncbi.nlm.nih.gov/). The coding regions for these known genes may be amplified and/or expressed using the techniques disclosed herein or as would be know to those of ordinary skill in the art. The expression of proteins, including enzymes, is generally known to those of skill in the art of molecular biology, for example, see Sambrook et al. (1989), incorporated herein by reference. The technique for preparing new monoclonal antibodies to a particular epitope to be detected as an analyte is quite straightforward, and may be readily carried out using techniques well known to those of skill in the art, as exemplified by the technique of Kohler and Milstein (1975), incorporated herein by reference.
In a preferred embodiment, the hydrogel is further defined as a hydrogel microsphere. In particularly preferred aspects of this embodiment, the hydrogel is fabricated from poly(ethylene glycol) diacrylate, sodium alginate (a.k.a. alginic acid) or a combination thereof. In other preferred aspects the hydrogel further comprises acrylic acid, allyl amine, styrene, allyl alcohol, acrylamide or a combination thereof. In a particularly preferred aspect, the hydrogel microsphere surrounds a liquid core. In certain aspects, the liquid core comprises water, alginic acid, or a co-polymer of poly ethylene glycol and poly isopropyl acrylamide, or a combination thereof. In certain aspects, it is preferred that the analyte sensitive compound or construct is contained or entrapped within said liquid core.
In other embodiments, the hydrogel undergoes a phase change to a solid state at a temperature from about 22xc2x0 C. to about 37xc2x0 C. This embodiment allows the hydrogel to remain in a fluid state until placed within a living organism, or other environment at or about 22xc2x0 C. or greater in temperature. This embodiment allows the creation of particularly shaped hydrogel structures in living organisms or environments at or about 22xc2x0 C. or greater in temperature. This embodiment is particularly useful to prepare the hydrogel materials at low temperatures for storage until use, or for molding the hydrogel into a desired shape.
In other embodiments, the composition is further defined or shaped as an at least one layer of hydrogel material. The hydrogel may further comprise a plurality of layers. In a preferred aspect, at least one layer contacts at least one successive layer to form a plurality of layers. A layer may be a net positively charged layer, a net negatively charged layer or an essentially neutrally charged layer. In a preferred aspect, the positively charged layer comprises an osmium derivative, a ruthenium derivative, a ferrocene derivative, a positively charged protein or a combination thereof. In another preferred aspect, the negatively charged layer comprises a negatively charged protein, such as for example, glucose oxidase. The positively or negatively charged protein or agent may provide the net positive or negative charge to the layer, respectively. Thus, for example, a layer that comprises a negatively charged protein may have a net negative charge because of its negatively charged protein content.
In certain aspects, the composition is shaped into a pattern, particularly to aid in the conveyance of information upon binding of an analyte to the composition, or to enhance sensitivity of analyte detection. The layer may be any shape, but recognizable symbols such as numbers, letters, geometric shapes, such as circles, squares, stars, and the like are preferred. In another aspect, the composition comprises a central body of hydrogel material. In another aspect, the composition comprises more than one body of hydrogel material. The more than one body of hydrogel material may be placed in multiple locations to form a pattern. In a particularly preferred aspect, the hydrogel is patterned as an array. The hydrogel matterial may be placed on a substrate or mold to aid its retention of shape or pattern. In certain preferred aspects, at least one arm of hydrogel material extends from central body of material. In a particularly preferred aspect, the least one arm comprises the analyte sensitive material. Such a configuration of projecting hydrogel material may enhance the speed and sensitivity of the analyte sensitive compound in detection of an analyte by aiding the ability of the material to contact the analyte.
The invention also provides a composition comprising an analyte sensitive detection system comprised within a hydrogel, wherein the hydrogel material includes but is not limited to, poly(ethylene glycol), poly(ethylene glycol)-co-anhydride, poly(ethylene glycol)-co-lactide, poly(ethylene glycol)-co-glycolide, poly(ethylene glycol)-co-orthoester, poly isopropylacrylamide, polyHEMA, polyacrylamide, sodium alginate or a combination thereof. The analyte sensitive detection system produces an optical or an electrochemical change upon contact with a selected analyte. In a preferred aspect, the hydrogel comprises polymerized poly isopropylacrylamide, polymerized poly(ethylene glycol) or sodium alginate.
In certain embodiments, the analyte sensitive detection system comprises an enzyme, such as glucose oxidase, galactose oxidase, cholesterol oxidase, cholesterase, lactate oxidase, lactate dehydrogenase or bilirubin oxidase, with glucose oxidase being preferred in certain aspects. Organophosphatase(s) are preferred enzymes to detect chemicals associated with chemical weapons. Paraxaon, sarin, tabun and samon are preferred organophosphate analytes, associated with chemical weapons, to be detected.
In preferred aspects of the invention, the analyte sensitive detection system produces a fluorescence change upon contact with the selected analyte. In these aspects, the analyte sensitive detection system or the hydrogel is in operable association with at least a first fluorescent label.
In certain aspects of the invention, the analyte sensitive detection system comprises at least a first component, the at least a first component described as an analyte binding component, wherein binding of an analyte to the analyte binding component alters the fluorescence of the analyte sensitive detection system. Depending on the particular system utilized, binding of the analyte to the analyte binding component can either increase or decrease the fluorescence of the analyte sensitive detection system.
In additional embodiments, the analyte sensitive detection system further comprises a second component that binds to the analyte binding component, wherein binding of the analyte to the analyte binding component decreases the binding of the analyte binding component and the second component. In certain aspects, the analyte competes with the second component for binding to the analyte binding component.
In particular embodiments, the analyte binding component comprises a first fluorophore conjugate and the second component comprises a second fluorophore conjugate. In preferred aspects, the fluorescence of the first fluorophore conjugate is quenched by the second fluorophore conjugate, exemplified by systems wherein the first fluorophore conjugate is FITC-dextran and the second fluorophore conjugate is TRITC-concanavalin A.
In particularly preferred aspects of the present invention, the compositions are formulated for implantation into an animal. In certain embodiments, the animal is a human.
The present invention also provides methods of detecting an analyte, comprising contacting a sample suspected of containing the analyte with an analyte sensitive detection system comprised within a hydrogel, wherein the hydrogel includes but is not limited to poly(ethylene glycol), poly(ethylene glycol)-co-anhydride, poly(ethylene glycol)-co-lactide, poly(ethylene glycol)-co-glycolide, poly(ethylene glycol)-co-orthoester, poly isopropylacrylamide, polyHEMA, polyacrylamide, sodium alginate or a combination thereof. An analyte that is detectable by the methods of the present invention may include, but is not limited to, a carbohydrate, a protein, a nucleic acid, a lipid, a gas or a chemical. Among the preferred analytes detected in the methods of the present invention are glucose, galactose, cholesterol, lactate, bilirubin, a blood gas, urea, creatinine, phosphate, myoglobin or a hormone, such as estrogen or progesterone. In particularly preferred aspects of the invention, the analyte detected is glucose. In other particularly preferred aspects, the analyte is an organophosphate. Organophosphatase(s) are preferred enzymes to detect chemicals associated with chemical weapons. Paraxaon, sarin, tabun and samon are preferred analytes, associated with chemical weapons, to be detected.
In certain aspects, the analyte sensitive detection system produces an electrochemical or optical change upon contact with the analyte. In a preferred aspect, the analyte sensitive detection system produces a fluorescence change upon contact with the analyte. In some aspects, the analyte sensitive detection system comprises at least a first component, the first component further described as an analyte binding component. In other aspects, the analyte sensitive detection system further comprises at least a second component. In additional aspects, the analyte binding component comprises at least a first fluorophore conjugate. In further aspects, the second component comprises at least a second fluorophore conjugate. In a preferred aspect, the fluorescence of the first fluorophore conjugate is quenched by the second fluorophore conjugate. In a particularly preferred aspect, the analyte is glucose, the first fluorophore conjugate is FITC-dextran and the second fluorophore conjugate is TRITC-concanavalin A.
In certain aspects, the sample suspected of containing the analyte is comprised within an animal. In this aspect, the analyte sensitive detection system may be formulated for implantation into an animal.
The invention further provides a method of using a smart tattoo, comprising implanting below the surface of the epidermis of the animal a smart tattoo comprising a hydrogel and an analyte detection compound, wherein the hydrogel comprises a polymerized material including but not limited to poly(ethylene glycol), poly(ethylene glycol)-co-anhydride, poly(ethylene glycol)-co-lactide, poly(ethylene glycol)-co-glycolide, poly(ethylene glycol)-co-orthoester, poly isopropylacrylamide, polyHEMA, polyacrylamide, sodium alginate or a combination thereof, and wherein the smart tattoo is implanted between about 0.05 mm and about 4 mm below the surface of the epidermis of the animal. As used herein certain embodiments, a xe2x80x9csmart tattooxe2x80x9d refers to a hydrogel and an analyte detection compound that may be implanted in the epidermis or dermis of an animal.
The epidermis may vary in thickness depending upon its location and the animal, but is generally up to about 1 mm thick in a human. When implanted in the epidermis, it is preferred that the tattoo is placed or implanted of from about 0.05 mm, about 0.06 mm, about 0.07 mm, about 0.08 mm, about 0.09 mm, about 0.1 mm, about 0.12 mm, about 0.14 mm, about 0.16 mm, about 0.18 mm, about 0.2 mm, about 0.22 mm, about 0.24 mm, about 0.26 mm, about 0.28 mm, about 0.3 mm, about 0.32 mm, about 0.34 mm, about 0.36 mm, about 0.38 mm, about 0.40 mm, about 0.42 mm, about 0.44 mm, about 0.46 mm, about 0.48 mm, about 0.50 mm, about 0.52 mm, about 0.54 mm, about 0.56 mm, about 0.58 mm, about 0.6 mm, about 0.62 mm, about 0.64 mm, about 0.66 mm, about 0.68 mm, about 0.7 mm, about 0.72 mm, about 0.74 mm, about 0.76 mm, about 0.78 mm, about 0.80 mm, about 0.82 mm, about 0.84 mm, about 0.86 mm, about 0.88 mm, about 0.90 mm, 0.92 mm, about 0.94 mm, about 0.96 mm, about 0.98 mm, to about 1 mm below the outer surface of the epidermis of an animal. In another preferred aspect, the smart tattoo is implanted between about 0.1 mm and about 0.25 mm below the surface of the epidermis of the animal. In a particularly preferred aspect, the smart tattoo is implanted about 0.15 mm below the surface of the epidermis of the animal. Preferred animals include sheep, goats, cats, dogs, birds, cows, horses or pigs. A particularly preferred animal is a human.
When implanted in the epidermis of an animal, the smart tattoo may exist only days or weeks before the cells containing or surrounding the tattoo are shed from the animal. In this embodiment, the tattoo will exist up to about 2 weeks before removal through natural replacement of epidermal layers.
In another embodiment, the tattoo is implanted in the dermis or dermal layers of an animal. The dermis may vary in thickness depending upon its location and the animal, but is generally from about 1 mm to about 4 mm thick in a human. The dermis is located beneath the epidermis, often generally beginning about 1 mm beneath the outer surface of the epidermis. The dermis does not actively shed, so that a tattoo may exist semi-permanently or permanently in an animal, i.e. remain in the dermis for months or years. Depending on the thickness of the epidermis and dermis, in certain embodiments, the tattoo may be implanted or placed in the dermis of from about 1 mm, about 1.1 mm, about 1.2 mm, about 1.3 mm, about 1.4 mm, about 1.5 mm, about 1.6 mm, about 1.7 mm, about 1.8 mm, about 1.9 mm, about 2 mm, about 2.1 mm, about 2.2 mm, about 2.3 mm, about 2.4 mm, about 2.5 mm, about 2.6 mm, about 2.7 mm, about 2.8 mm, about 2.9 mm, about 3 mm, about 3.1 mm, about 3.2 mm, about 3.3 mm, about 3.4 mm, about 3.5 mm, about 3.6 mm, about 3.7 mm, about 3.8 mm, about 3.9 mm, about 4 mm, about 4.1 mm, about 4.2 mm, about 4.3 mm, about 4.4 mm, about 4.5 mm, about 4.6 mm, about 4.7 mm, about 4.8 mm, about 4.9 mm, to about 5 mm beneath the outer surface of the epidermis. In certain preferred embodiments the tattoo would be implanted of from about 1 mm to about 5 mm beneath the surface of the epidermis, with 2 mm to about 3 mm being particularly preferred.
When implanted in tissues, the composition may be taken into a cell or remain external to a cell. The particle size of the composition, and its size ratio to that of the size of an adjacent cell will determine whether the composition is taken into a cell or remains external to the cell. In certain embodiments wherein the composition is implanted adjacent to epidermal and dermal cells, compositions of an average particle size up to about 10 microns in diameter or so may be taken into cells. In certain other embodiments, a composition of an average particle size of from about 0.5 microns, about 1 microns, about 2 microns, about 3 microns, about 4 microns, about 5 microns, about 6 microns, about 7 microns, about 8 microns, about 9 microns, to about 10 microns or so may be taken into cells. In certain embodiments, it is preferred that the composition is taken into a cell to measure the intracellular concentration of an analyte. For example, intracellular glucose levels may vary more relative to plasma glucose concentrations in diabetics. Detection of low intracellular glucose levels may aid in monitoring changes in glucose in diabetes or the effectiveness of medications.
In certain other embodiments, it is preferred that the composition remains external to the cells of the tissue that the composition is implanted. A larger average particle size is preferred for the composition in this embodiment, to prevent the composition""s uptake by cells. In certain embodiments for epidermal or dermal cells, a composition greater than about 10 microns in average particle size diameter is preferred. In certain embodiments, a composition of about 10 microns, about 11 microns, about 12 microns, about 13 microns, about 14 microns, about 15 microns, about 16 microns, about 17 microns, about 18 microns, about 19 microns, about 20 microns, about 22 microns, about 24 microns, about 26 microns, about 28 microns, about 30 microns, about 35 microns, about 40 microns, about 45 microns, about 50 microns, about 55 microns, about 60 microns, about 65 microns, about 70 microns, about 75 microns, about 80 microns, about 85 microns, about 90 microns, about 95 microns, about 100 microns, about 110 microns, about 120 microns, about 130 microns, about 140 microns, about 150 microns, about 160 microns, about 170 microns, about 180 microns, about 190 microns, about 200 microns, about 225 microns, about 250 microns, about 275 microns, about 300 microns, about 350 microns, about 400 microns, about 450 microns, about 500 microns, about 550 microns, about 600 microns, about 650 microns, about 700 microns, about 800 microns, about 850 microns, about 900 microns, about 950 microns, about 1 mm, about 1.1 mm, about 1.2 mm, about 1.3 mm,
about 1.4 mm, about 1.5 mm, about 1.6 mm, about 1.7 mm, about 1.8 mm, about 1.9 mm, to about 2 mm or more average particle size is preferred.
Of course, different analyte detection compounds may be combined with different particle sizes in various combinations. In one embodiment, particles of one average sized diameter may detect a different analyte than another composition with a different average particle size diameter. Alternatively, detection compounds that detect the same analyte may comprise compositions of functionally similar size. In certain embodiments, a different optical and/or electrochemical detection agent comprises the analyte detection compound. Thus, by combining various particle sizes with different analyte detection compounds, composition with different detection properties may be created. For example, fluorescence at one wavelength in a composition with an average particle size less than about 10 microns may denote the presence of a certain concentration of an analyte in the intracellular spaces, while fluorescence at a different wavelength in a composition with an average particle size greater than about 10 microns may denote the concentration and the intercellular spaces of a tissue. In another example, the composition may comprise more than one analyte detection compounds to detect different concentrations of an analyte, and/or different analytes.
In certain aspects, the smart tattoo may detect an analyte that may include, but is not limited to a carbohydrate, a protein, a nucleic acid, a lipid or a gas. In preferred aspects, the analyte is glucose, cholesterol, lactate, bilirubin, a blood gas, urea, creatinine, phosphate, myoglobin or a hormone. In a particularly preferred aspect, the analyte is glucose. In a preferred aspect, the smart tattoo produces an electrochemical or an optical change upon contact with the analyte. A preferred optical change is a fluorescence change upon contact with the analyte.
In certain aspects, the smart tattoo comprises at least a first component, the first component further described as an analyte binding component In additional aspects, the smart tattoo further comprises at least a second component. In a preferred aspect, the analyte binding component comprises at least a first fluorophore conjugate. In another preferred aspect, the second component comprises at least a second fluorophore conjugate. In certain aspects, the fluorescence of the first fluorophore conjugate is quenched by the second fluorophore conjugate. In a particularly preferred aspect, the analyte is glucose, the first fluorophore conjugate is FITC-dextran and the second fluorophore conjugate is TRITC-concanavalin A.
The efficiency of detection of an optical change in an analyte is dependent upon the wavelength of light used to visualize the composition. Shorter wavelengths, such as the near UV to blue part of the spectrum, i.e. about 350 nm to about 450 nm, are preferred to detect optical changes in a composition implanted up to about 0.4 mm beneath the surface of the epidermis. Longer wavelengths may penetrate deeper into tissue, and wavelengths of the yellow-orange-red-near infrared part of the spectrum, i.e. greater than about 450 nm to about 2 mm or greater, are preferred to detect optical changes in a composition implanted up to about 2 or about 3 mm beneath the surface of the epidermis. In certain embodiments, compositions may be created that are normally be invisible to the naked eye without illumination with a light source, detection with one or more electrodes, or fluorescence or phosphorescence of the composition.
In certain embodiments, the analyte sensitive detection system produces an electrochemical change upon contact with the analyte, while in other embodiments, the analyte sensitive detection system produces an optical change, for example a fluorescence change, upon contact with the analyte. In a preferred embodiment, the hydrogel wherein the analyte sensitive detection system is contained or attached is in contact with a substrate, such as for example, an electrode.
In particular aspects, the analyte sensitive detection system comprises at least a first component, the first component further described as an analyte binding component, while in additional aspects, the analyte sensitive detection system further comprises at least a second component. In certain embodiments, the analyte binding component comprises at least a first fluorophore conjugate. In other aspects, the second component comprises at least a second fluorophore conjugate. In particularly preferred aspects of the invention, the fluorescence of the first fluorophore conjugate is quenched by the second fluorophore conjugate, exemplified by embodiments wherein the analyte is glucose, the first fluorophore conjugate is FITC-dextran and the second fluorophore conjugate is TRITC-concanavalin A.
In preferred embodiments of the present invention, the sample suspected of containing the analyte is comprised within an animal. In further preferred embodiments, the analyte sensitive detection system is formulated for implantation into an animal. In certain embodiments, the animal is a human.
Following long-standing patent law convention, the word xe2x80x9caxe2x80x9d and xe2x80x9canxe2x80x9d mean xe2x80x9cone or morexe2x80x9d in this specification, including the claims.